Replacing or supplementing fractured, damaged, or degenerated mammalian skeletal bone with prosthetic implants made of biocompatible materials is commonplace in the medical arts. Most often, implant devices are intended to become permanently integrated into the skeletal structure. Unfortunately, permanent prosthetic attachment to bone is rare. Factors that influence long-term implant viability include material type used, bone fixation method, implant location, surgical skill, patient age, weight and medical condition. A plethora of devices have been constructed attempting to optimize these variables involved in producing an increase in bone fusion.
Common materials used in prosthetic devices include ceramics, polymers and metals. Currently, metallic materials afford the best mechanical properties and biocompatibility necessary for use as skeletal prosthetic implants. Frequently used metals include, titanium and titanium alloy, stainless steel, gold, cobalt-chromium alloys, tungsten, tantalum, as well as, similar alloys. Titanium is popular in the implant field because of its superior corrosion resistance, biocompatibility, physical and mechanical properties compared to other metals. The dramatic increase over the last decade of titanium material presentations in neurosurgical, orthopedic and dental surgery attests to its acceptance as a prosthetic material. Titanium presentations vary mostly in shape and surface type, which influence the implant's ability to support load and attach to bone.
A significant drawback to titanium implants is the tendency to loosen over time. There are three typical prevailing methods for securing metal prosthetic devices in the human body: press-fitting the device in bone, cementing them to an adjoining bone with a methacrylate-type adhesives, or affixing in place with screws. All methods require a high degree of surgical skill. For example, a press-fitted implant must be placed into surgically prepared bone so that optimal metal to bone surface area is achieved. Patient bone geometry significantly influences the success of press-fitted implants and can limit their usefulness as well as longevity. Similar problems occur with cemented implants; furthermore, the cement itself is prone to stress fractures and is not bio-absorbable. Therefore, all methods are associated to varying degrees with cell lysis next to the implant surface with concomitant fibrotic tissue formation, prosthetic loosening, and ultimate failure of the device.
Currently, methods are being developed that produce osteointegration of bone to metal obviating the need for bone cements. Osteointegration is defined as bone growth directly adjacent to an implant without an intermediate fibrotic tissue layer. This type of biologic fixation avoids many complications associated with adhesives and theoretically would result in the strongest possible implant-to-bone bond. One common method is to roughen a metal surface creating a micro or macro-porous structure through which bone may attach or grow. Several implant device designs have been created attempting to produce a textured metal surface that will allow direct bone attachment. Some of these devices are found in the following U.S. Pat. Nos.: 3,894,297; 3,905,777; 3,906,550; 4,064,567; 4,199,824; 4,261,063; 4,430,761; 4,479,271; 4,530,116; 4,535,487; 4,536,894; 4,549,319; 4,570,271; 4,589,883; 4,608,053; 4,636,219; 5,018,285; 5,344,654; 5,373,621; 5,609,635; and 5,658,333.
Metallic implant surfaces are also commonly coated with micro-porous ceramics such as hydroxyapatite (HA) or beta-tricalcium phosphate (TCP) (see U.S. Pat. Nos. 4,309,488; 4,145,764; 4,483,678; 4,960,646; 4,846,837). The former treatment is more common because calcium-phosphate salts tend to be absorbed, in vitro, and thus loose their effectiveness. The HA coatings increase the mean interface strength of titanium implants as compared to uncoated implants (see Cook et al., Clin. Ortho. Rel. Res., 232, p. 225, 1988). In addition, clinical trials in patients with hip prosthesis have demonstrated rapid bone growth on prosthetic devices and increased osteointegration of titanium alloy implants when coated with HA (see Sakkers et. al., J. Biomed. Mater. Res., 26, p. 265, 1997). The HA ceramic coatings can be applied with a plasma spray machine or by sintering (see U.S. Pat. No. 4,960,646). In addition, the HA coating can be applied by soaking the implant in an alkali solution that contains calcium and phosphorous and then heated to deposit a film of hyroxylapetite (see U.S. Pat. No. 5,609,633). Optimal HA coating thickness ranges from 50-100 microns (see Thomas, Orthopedics, 17, p. 267-278, 1994). If coated too thick the interface between the HA and bone becomes brittle. Despite the higher success rate of prosthetic devices coated with HA as compared to earlier implantation methods, failure over time still occurs. Again, proper integration requires that the surgeon create an exact implant fit into bone allowing the metal and bone surfaces to have maximum contact. Also, fibrotic tissue formation develops in some cases regardless of coating type.
Recent research describes the use of osteoinductive proteins to produce prosthetic osteointegration as well as increase the rate of bone formation next to implant surface (for example see Cole et. al., Clin. Ortho. Rel. Res., 345, p.219-228, 1997). Osteoinductive proteins are secreted signaling molecules that stimulate new bone production. These proteins include, PDGF, IGF-I, IGF-II, FGF, TGF-.beta. and associated family members. The ability of these proteins to enhance osteointegration of metallic implants suggests that implants coated with these proteins may attach to bone more efficiently.
The most effective bone formation-inducing factors are the bone morphogenetic proteins (BMPs). The BMPs, a TGF .beta. super-family subset, share, along with the other members of its subgroup, strong sequence homology and conserved carboxyl-terminus cysteine residues. Over 15 different BMPs have been identified. Most members of this TGF-62 subfamily stimulate the cascade of events that lead to new bone formation (see U.S. Pat. Nos. 5,013,649; 5,635,373; 5,652,118; and 5,714,589, reviewed in J. Bone Min. Res., 1993, v8, suppl-2, p.s565-s572). These processes include stimulating mesenchymal cell migration, osteoconductive matrix disposition, osteoprogenitor cell proliferation and differentiation into bone producing cells. Effort, therefore, has focused on BMP proteins because of their central role in bone growth and their known ability to produce bone growth next to titanium implants (see Cole et. al., Clin. Ortho. Rel. Res., 345, p.219-228, 1997). One such method claims achievement of a strong bond between existing bone and the prosthesis by coating the prosthetic device with an osteogenic protein (see U.S. Pat. No. 5,344,654).
In addition to osteoinductive proteins, osteoconductive factors may aid in bone formation (see U.S. Pat. No. 5,707,962). One experienced in the art realizes that osteoconductive factors are those that create a favorable environment for new bone growth, most commonly by providing a scaffold for bone ingrowth. The clearest example of an osteoconductive factor is the extracellular matrix protein, collagen. Other factors that can be considered osteoconductive include nutrients, anti-microbial and anti-inflammatory agents, as well as blood-clotting factors. In addition to these factors, reducing bone absorption by inhibiting osteoclast activity with bisphosphonate may also aid in implant success (see U.S. Pat. No. 5,733,564).
Bone morphogenetic protein-molecule presentation to skeletal tissue is critical for producing desired bone formation next to an implant device. Many matrix systems have been developed to contain and then steadily release bioactive peptides as the matrix degrades. Organic polymers such as polylactides, polyglycolides, polyanhydrides, and polyorthoesters, which readily hydrolyze in the body into inert monomers, have been used as matrixes (see U.S. Pat. Nos.: 4,563,489; 5,629,09; and 4,526,909). The efficiency of BMP-release from polymer matrixes depends on the matrixes resorbtion rate, density, and pore size. Monomer type and their relative ratios in the matrix influence these characteristics. Polylactic and polyglycolic acid copolymers, BMP sequestering agents, and osteoinductive factors provide the necessary qualities for a BMP delivery system (see U.S. Pat. No. 5,597,897). Alginate, poly(ethylene glycol), polyoxyethylene oxide, carboxyvinyl polymer, and poly (vinyl alcohol) are additional polymer examples that optimize BMP-bone-growth-induction by temporally sequestering the growth factors (see U.S. Pat. No. 5,597,897).
Non-synthetic matrix proteins like collagen, glycosaminoglycans, and hyaluronic acid, which are enzymatically digested in the body, have also been used to deliver BMPs to bone areas (see U.S. Pat. Nos.: 4,394,320; 4,472,840; 5,366,509; 5,606,019; 5,645,591; and 5,683,459). In human bone, Collagen serves as the natural carrier for BMPs and as an osteoconductive scaffold for bone formation. Demineralized bone in which the main components are collagen and BMPs has been used successfully as a bone graft material (see U.S. Pat. No. 5,236,456). The natural, or synthetic, polymer matrix systems described herein are moldable and release BMPs in the required fashion; however, used alone these polymers serve only as a scaffold for new bone formation. For example, U.S. Pat. Nos. 5,683,459 and 5,366,509 describe an apparatus, useful for bone graft substitute, composed of BMPs injected into a porous polylactide and hyaluronic acid meshwork. Furthermore, an osteogenic device capable of inducing endochondral bone formation when implanted in the mammalian body has been disclosed (see U.S. Pat. No. 5,645,591); this device is composed of an osteogenic protein dispersed within a porous collagen and glycosaminoglycan matrix. These types of devices were designed as an alternative bone graft material to replace the more invasive autograft procedures currently used. These devices by themselves would not work well as joint prosthesis due to their brittle nature and constant joint movement preventing bone formation into the device.
Proper implant load distribution is yet another characteristic important for correct prosthetic function. This issue prompted the development of a variety of devices that attempt to distribute the weight bearing load of the prosthetic implant or produce a direct bone-implant bond. For example, U.S. Pat. No. 5,639,237 describes an endosseous dental implant having a dimpled surface texture for use in crania-facial bones reconstruction. The indented surface increases contact area for bone proliferation, thereby enhancing the dental implant mechanical fixation or anchoring strength as compared to ordinary dental implants having a similar geometry. A similar prosthetic device manufacturing method for securing implant into human bone is described in U.S. Pat. No. 5,360,446. A description of an altogether different technique for enhancing bone density adjacent to the implant is in U.S. Pat. No. 5,344,457. This reference teaches that loading stress can be effectively transferred from a dental implant to surrounding bone through a tapered body shaped implant. Yet another technique is described in U.S. Pat. No. 5,458,653. This reference describes a prosthetic device coated with a bioabsorbable polymer in specific implant regions to, theoretically, better distribute the load placed upon it. Many other endosseous dental implants with shapes attempting to distribute load including helical wires, tripods, screws and hollow baskets have also been used. The clinical success of all these implant types is dependent on placement site, implant fit and the extent of fibrous tissue formation around the implant preventing direct bone contact.
Further complications arise when placing a prosthetic implant in skeletal areas that cannot support large functional loads and sheer stresses. Crania-facial implants, which are commonly used in the reconstruction or replacement of single teeth, are particularly prone to failure from stress. These prosthetic failures are primarily due to the inability of cancellous bone to support implant load. Unlike smooth, densely packed cortical bone, cancellous bone is porous and has an asymmetric sponge-like structure. In addition, small bones of the hand, elbow and feet that do not have thick cortical walls are prone to implant failure. Implants in these areas often fail due to excessive movement and a lack of supporting bone structure. For example, a prosthetic joint replacement device, when used in small hand bones, will often become loose and erode the surrounding bone due to lack of cortical structure. Methods, therefore, have been devised to augment or support porous or less dense bone.
One method for augmenting or supporting porous or less dense bone is provided in U.S. Pat. Nos. 4,693,721 and 5,030,233, wherein a biocompatible porous titanium-wire-mesh material is described for use in bone repair or replacement. Presumably, the porous mesh, when implanted, would allow bone ingrowth and distribute stress load while reinforcing areas of low density bone. Furthermore, according to U.S. Pat. Nos. 3,906,550, 4,660,755, 4,829,152 and 4,923,513, a porous titanium matrix can be welded onto a solid titanium prosthetic implant. Presumably, these methods would produce a broader load distribution across the titanium mesh surface, increase prosthetic implant surface area and load distribution, as well as reinforce bone areas that lack density. However, the fibrillar mesh of these described devices are cemented into place. In addition, these devices require that bone grow in-between the titanium-mesh is unaided by stimulating proteins. This bone growth is speculative and may or may not occur depending on the strength of the implantation site, bone health in the area treated, and the distance that the bone has to grow. In Michelson, U.S. Pat. No. 5,609,635, a method is described for the design of a spinal fusion device comprised of wire mesh infused with osteoinductive molecules. This device is intended solely for use in spinal fusions and is not designed for use with other prosthetic implants intended for use in other body areas. It is also not designed to be attached to orthopedic implants.
Despite the plethora of prior art approaches to securing an implanted structure into mammalian bone, there is a need in the medical and dental arts for improving the strength and integrity of the bone that surrounds and attaches to a prosthetic implant device. Furthermore, there is a need in these arts for a device that both produces osteoinductive-protein-induced bone formation between metal fibers and increases structural bone integrity, as well as the bone to implant contact area. As will be seen, the present invention provides a method and a structure for increasing strength and distribution of load bearing areas of the bone surrounding a prosthesis. In addition, the present invention provides a novel way to augment both endo and exo bone formation for a variety of applications.